The numbers in brackets below refer to publications listed in the Appendix, the teachings of which are incorporated herein by reference.
When viewed from an engineering perspective the human skeleton is a marvelous engineering structure. For the long bones, for example, the cartilage-covered epiphyseal surfaces of articulating joints cannot carry heavy stresses [1]. So the joints are enlarged to reduce that stress and provide smooth articulating surfaces with extremely low friction. These bearings are supported by a thin layer of dense bone supported by trabecular bone, with the trabecular bone aligned in accordance with Wolff's law to convey the low-stress bearing forces to the dense, high-stress cortical bone of the diaphysis [2]. This remarkable structure is even more remarkable because, during growth, the dense ring of diaphysis bone just below the growth plate is continuously elongated, with cartilage calcifying below the growth plate and chondrocytes building cartilage above, all without loss of strength as the process operates to elongate the bones [3]. In this process the cell differentiation, and the vascularization and ossification mechanisms, operate continuously to produce the dense bone of the diaphysis, and still maintain the large-area bearing surfaces. The stiffness of the diaphysis, with its dense cortical bone forming a tube, is much greater than that of a solid rod of the same mass. The exterior of the epiphysis is a structural membrane. From a mechanical point of view this is an example of a marvelous engineering structure.
The physiological process mentioned above can also be approached from an engineering point of view. It can be seen that, in each part of the structure there is a supply of cells, nutrients, enzymes, and chemicals to provide for growth, repair, and remodeling. Using the long bone joint as an example, provision for cellular activity is essential for growth, repair, and remodeling. For growth to take place there must be a steady supply of the necessary nutrients. Mechanisms for controlling that supply are extraordinarily complex [4]. For example, at the growth ring, formation of cartilage is inhibited by its own growth, causing differentiation of chondrocytes into osteoblasts. The perichondrium becomes a periosteum. The chondrocytes hypertrophy and die, producing a collagenous matrix containing cavities left by the empty chondrocyte lacunae. Osteoblasts line the cartilage and synthesize osteoid. The cartilage begins to calcify. This mineralization reduces diffusion of nutrients and the osseous tissue contracts, producing dense osseous tissue of smaller diameter, that of the diaphysis. The mechanisms are not understood. For example, the mechanism of precipitation of hydroxyapatite is unknown [5]. Transport of calcium and phosphate ions proceed separately. Nucleation may take place homogeneously in matrix vesicles, or heterogeneously at collagen fibril voids [6]. Regardless of the mechanisms and the complexities, however, this can be considered analogous to an engineering system, something like the logistics of battle support. The necessary nutrients, mechanisms, and other factors must be provided in a timely manner for the tissue to remain healthy and strong. This can include electrical transport, the effects of stress, and many other factors. In order to have a successful prosthesis, then, all aspects of skeletal engineering must be considered. Unfortunately this has not been done.
Orthopedic research can be divided into two areas, clinical and academic. Clinical research includes the improvements and modifications of existing prostheses and surgical procedures to improve patient care. Outstanding surgeons often work with prosthesis manufacturers to modify the existing practice in attempts to improve the performance [7]. Recently this has included the introduction of beads, mesh, and other structures on the surface of metallic implants to provide for tissue ingrowth and assist in stabilization [8]. Unfortunately, a roughened implant can also cause tissue irritation and lead to failure. New ideas, such as the substitution of pure titanium for the 6Al4V alloy, have been tested in the clinic. In the dental area, for example, two tooth-root prostheses made of carbon have been introduced into clinical practice and subsequently withdrawn because of failure [9, 10]. Thus the industry often is doing its research in the clinic.
Academic research is conducted through government laboratories, private laboratories, and academia. Here the research is highly specialized. Each individual specialty is dominated by the theory and practice of that specialty. For example, new materials are often evaluated on the basis of specific cellular responses to those materials regardless of whether the material is successful [11]. For example, carbon/Teflon mesh has been tested as an orthopedic material despite the fact that it is easily crushed in one's fingers [12]. Also, only after other countries adopted the idea of osseous integration was the Branemark titanium implant for tooth roots accepted in the United States. Before that time osseous integration was not considered grounds for acceptance.
The special nature of academic research has been allowed to dictate the research results. For example, bone and tissue interactions are often modeled by finite element methods to determine stress distributions in the prosthesis and in the bone [13]. But bone is complex. Almost all of these efforts have modeled bone as a continuum. In fact, bone is not isotropic and not homogeneous. Even today most finite element analyses are conducted assuming cortical bone has one set of homogeneous properties and cancellous bone has another set of homogeneous properties [14]. Actual bone varies tremendously, and the structure of the bone must be considered. When this is not done the results may be wrong and misleading. An even more serious criticism of this research is that the bone changes in response to the prosthesis. Where stresses are very high, bone is resorped, distributing the stress and changing the geometry. This is not modeled, so the model is inaccurate as soon as the implant is in place.
The political desires of research funding also strongly affect academic research. There is a large funded effort to study orthopedic materials in vitro, to avoid in vivo studies. Programs in cell attachment and other areas are funded despite the fact that the correlation of in vivo with in vitro has not been established [15].
The research method of academia is to divide and conquer; to study one variable while keeping all others constant. This advances fundamental knowledge but does not consider the interaction of various factors. Real orthopedic implants require the simultaneous application of chemistry, basic biological science, physiology, anatomy, materials science, stress analysis, and systems analysis. The bottom line is successful orthopedic performance. This requires synthesis from all pertinent areas, identification of the important and the trivial factors, experimentation, and decision making. This is the essence of engineering design.
Each year in the United States about 250,000 total hips replacement surgeries are performed. And about 25,000 hip prosthesis replacements are performed. The number of replacements is expected to climb because the life expectancy of a hip is 5 to 15 years. Patients are living longer. There is a need for improved hip prostheses.
Tissue response is critical to the success of an implant. It is the tissue response that determines the life of the implant because the hard tissue is continuously remodeling. Only if the tissue continues to support the implant can it be successful. The tissue response depends, in part, on the material from which it is made.
There are three classes of materials: metals, organic materials, and ceramics. The wrong material invokes classical undesirable tissue response. This includes inflammation, the presence of macrophages, a fluid-filled capsule, and resorption of surrounding tissue [16]. Most materials in all three classes invoke this response if they have any solubility in the tissue fluids. This puts serious limitations on the materials that can be considered. All the structural components of existing prostheses are selected to be bioinert [17]. These have minimum solubility. Metals such as titanium 316-L stainless steel, Al6V4 titanium, and cobalt-chrome alloys all are inert; although, especially in wear situations, some solubility and tissue reaction does occur. The organic polymers such as very high molecular weight (VHMW) polyethylene and polymethylmethracralate also are chosen for their minimum tissue response. Ceramic components such as alumina and zirconia are also bioiner. As such, all the inert materials are foreign bodies and are walled off by a thin fibrous capsule. The better materials have thinner capsules. However, there is not a direct bond of osseous tissue to bone. Such implants are only successful if the remaining tissue continues to support it. This will be time dependent because any movement of the prosthesis will increase the thickness of the fibrous capsule and because of the tissue degeneration described above, leading to more movement and progressive failure. The undesirable effects of electrochemical cells for electronic conductors such as metals and the degeneration of organics such as methacrylates and polyethylene would make the ceramic materials more attractive. Many ceramic compounds are highly insoluble and inert in a physiological environment, and most are electrical insulators. However, the ceramics are brittle. This limitation is discussed later.
Materials that are not inert but are not walled off by a foreign-body capsule have especially desirable tissue response. The only known materials of this nature are calcium phosphates. The ions released by the calcium phosphates, Ca.sup.2+ and (PO.sub.4).sup.3- are also present in bone. The natural mineral in bone is impure hydroxyapatite, Ca.sub.10 (PO.sub.4).sub.6 (OH).sub.2, which contains water. Ceramic processes that require firing at high temperatures remove all or part of the water in hydroxyapatite. The residue may be oxyapatite of the same crystalline structure, or it may disproportionate into Ca.sub.3 (PO.sub.4).sub.2 and Ca.sub.2 P.sub.2 O.sub.7. Both are only slightly soluble and hydrolyze on the surface to hydroxyapatite. The tissue response for calcium-to-phosphorous ratios between 1.5 [Ca.sub.3 (PO.sub.4).sub.2 ] and 1.67 [hydroxyapatite] is known to be extremely compatible with hard tissue. Their performance is often call osteoconductive to distinguish it from osteoinductive, the latter being the production of osseous tissue in soft tissue sites. The lack of a fibrous capsule and the ability of bone to bond to the calcium phosphates makes them very interesting for prosthesis applications. Tissue response is critical, and if calcium phosphates can be used to achieve a bond between the implant and the hard tissue, they make long life a possibility. It is the tissue life that is so important. Therefore, to obtain improvement in tissue response to the current metals and alloys, the use of calcium phosphates is the only choice known at the present time.
An obvious approach to utilizing calcium phosphates, where they do not have the required strength, is to incorporate them into a composite. Combinations with plastics, metals, and ceramics can be considered. Composites based on plastics were rejected because it was difficult to expose the calcium phosphate when embedded in plastics. The limited rigidity was also a detriment. Composites based upon carbon were rejected for the same reasons, although fiber reinforcement was a desirable possibility. Composites based on metals were rejected because of reactions with the metals, including phosphorous embrittlement. High-temperature processing of metals alters the physical properties and requires reducing atmospheres, whereas the phosphates require oxidizing atmospheres. Therefore, ceramic-ceramic composites appear to be more promising.
Not many ceramic materials have no reaction with calcium phosphates at processing temperatures (up to 1500.degree. C.). Based on physical chemistry and crystallographic considerations, magnesium aluminate spinel (MgAl.sub.2 O.sub.4) was selected as a suitable second phase to use with a calcium phosphate [25, 26]. The ionic size of Mg is too small for extensive solid solution in calcium phosphates. The aluminum ion is too large to proxy for phosphorous. Spinel is an extremely stable compound, as are tricalcium phosphate and calcium pyrophosphate. Spinel is known to be an excellent, inert ceramic in its tissue response. If the calcium phosphate is to control tissue response, it must be distributed on a very fine scale. If it is not to cause large flaws, it also must be very fine grained. The calcium phosphate should not be easily leached away, so the calcium phosphate phase should be interconnected, not isolated. For this to occur the calcium phosphate phase should be at least 25 vol %. Fifty percent would be better.
Hip replacements are indicated when the femoral stem breaks off or when degenerative arthritis has made the joint too painful for function. Most of the procedures are performed for arthritis relief. The present design requires the removal of the ball and the stem of the femur. Removal of strong healthy bone to accommodate a femoral stem is undesirable and can be avoided with the design disclosed here. Excessive reaming of the acetabulum should also be avoided.
Part of the reason for attempting replacement of existing polyethylene and metal femoral components with inert hard ceramics, such as alumina, is to avoid wear and wear debris. Reduction in friction helps in this objective. However, the existing designs are only slightly modified from the conventional metal and polyethylene design, are subject to brittle failure (fracture of the stem, cup, etc.), are difficult to install and are expensive. The ball and stem must be machined in a very complex shape. And the balls and cups are the same shape and size as the polyethylene component. The size of the ball depends on the stresses that can be applied to the components.
All orthopedic implants currently in use rely on mechanical fixation, whether metal, ceramic, or plastic. Examples include: acetabular components and femoral components of a total hip, where screws and posts are used for fixation; knees where the femoral component is held with screws and the tibial component is held with posts and screws. All such implants are walled off by the foreign body response. A fibrous capsule wall separates the implant from the hard issue. Loosening under stress as the fibrous capsule deforms is progressive, because motion prevents repair and the capsule becomes thicker, which causes further loosening, etc. This is the most common mode of failure, usually after about five years.
The foreign body response is defined as the separation or walling-off of a foreign body such as an implant with a soft, fibrous, collagenous layer that prevents the hydroxyapatite of the bone from touching the implant. The fibrous capsule becomes thicker if there is motion between the implant and the bone. The thicker capsule allows more motion, which causes a still thicker layer. This is progressive and is the most common mechanism of failure for implants using mechanical fixation. Bone screws and other anchoring devices suffer from this encapsulation and is the reason a new method of fixation is needed.
The major cause of replacement is loosening under stress. Attempts to reduce this defect include hydroxyapatite coatings, beads or mesh for tissue ingrowth and modification of the alloys. Titanium and its alloys, 316-L stainless steel and cobalt chromium alloys are used for the femoral component. Both the acetabular and the femoral component are often cemented in place with polymethacrylate cement. The cement then has the tissue contact.
The cement is inserted under pressure before the prosthesis component is inserted. This ensures the cement fills the space between the prosthesis and the tissue. When the cement polymerizes the prosthesis is fixed, and forces applied to the prosthesis component are transmitted through the cement to the bone. However the cement, inserted under pressure into trabecular bone, penetrates and displaces the soft tissue in the trabeculae, effectively shutting off the blood and nutrient avenues for repair of the trabecular walls. Therefore, the area of heathy tissue contact is at the periphery of the cement. Degeneration within the cement cannot be repaired, and the bone at the periphery is in contact with whatever fracturing and internal debris produced in the zone adjacent to the periphery. This can cause the bone at the periphery also to degenerate.
The major cause for loosening is the foreign-body response of the tissue. All the plastics, conventional ceramics, and metals are walled-off with a fibrous capsule by the body's foreign-body response mechanism. If roughness or micro-motion occurs the capsule gets thicker, leading to more motion, more thickness and progressing to failure. Although the materials are chosen to be inert, to have a very thin capsule, this is the most common mechanism of failure.
Exiting total hip prostheses have a dense polyethylene cup and a metal femoral ball-and-stem design. Attempts are being made, especially in Europe, to replace both components with ceramic (alumina or zirconia.) In the United States they have not generally been accepted in clinical practice. Other than infection there are three principal causes of failure. One is the general deterioration by osteoporosis, a degenerative condition that is difficult to prevent. The second is wear of the polyethylene or metal components, that produces debris that causes tissue damage and dysfunction. The third is progressive loosening of the acetabular or femoral component, that leads to pain, dysfunction and failure.
When a tensile strain is applied to a natural hip joint the ligaments and muscles inhibit dislocation of the joint. One of the problems of the implants currently used is frequent dislocation.
The orthopedic surgeon is often faced with clinical situations in which bone has been removed, shattered or missing. Cancer surgery, accident trauma and genetic defects are frequently causes. Whenever this happens, and the remaining bone cannot be re-joined and stabilized in a satisfactory way, some sort of bone grafting is necessary.
The need for bone grafts is the result of natural limitations for bone to repair itself. In the event of fractured or missing bone the remaining bone must be stabilized in its natural position by internal or external fixing methods, such as internal or external bone plates and plaster casts. This keeps the bone from moving so that the physiology of bone repair can occur; typically hematoma, fibrous callous, mineralized callous and remodeling. However, if much fragmentation has occurred or if too much bone is missing the bone does not reunite even if the remaining bone is stabilized. This typical non-union problem is one reason for considering bone grafts. In some species, such as dogs and cats, for example, it is know that fractures of the diaphysis of long bones result in non-union whenever the length of missing bone is more than 11/2 times the external diameter of the bone. Thus a method could be tested by grafting a longer defect in dogs and cats.
The best bone grafting material is natural bone, (autogenous bone) from somewhere else in the patient. This is used where practical. Unfortunately, it is often impractical because of the unavailability of suitable bone in the patient. Even when it can be obtained from the patient there is still the trauma of removal of bone from somewhere else in the body, the danger of fracture at the removal site, and the danger a second surgical procedure. Therefore, a source of bone grafts from some other source is essential in many clinical situations.
If autogenous bone is not available the surgeon often uses bone from the same species but a different individual in that species (allograft bone). This causes problems with the immune system and brings with it danger of infection. The success rate is lower than that of autogenous bone. Because of these risks a bone grafting method has long been sought for these difficult clinical situations. Much research has been conducted to find new materials to replace natural bone. This includes the full spectrum of biocompatible metals, organics (plastics) and ceramics. None of these has been successful because, in general, the inert materials chosen for the purpose are either walled off by the foreign body mechanism of the host (a fibrous capsule around the implant) or cause serious physiological responses when friction or chemical reactions produce particulates of the implant materials. Both causes are major sources of loosening of implants, pain and failure.
Many different devices and geometries of bone grafting have been attempted and many patents have been issued. Yet there is no standard material or method that can satisfy this common surgical problem. This is true because the problem is complex. It requires both material that is compatible with bone healing, restoration and remodeling mechanisms, and a way to utilize those mechanisms to repair the bone to its natural load-bearing state.
Those concerned with these and other problems recognize the need for an improved method of producing restructured bone and causing bone to bond to an implant.